Positron Emission Tomography (“PET”) is an imaging technique that can be used to develop two-dimensional and three-dimensional tomographic images of a distribution of positron-emitting isotopes within a subject, such as a human patient undergoing medical imaging in which the resulting PET images provide a visual depiction of tissue differences within the subject. PET procedures typically involve introducing one or more radiolabeled pharmaceutical tracer compounds into the subject, usually through injection or inhalation. As the radioisotope(s) incorporated in the pharmaceutical tracer compound decays, it releases positrons. These positrons collide with surrounding matter before combining with an electron in an annihilation event that destroys the positron and electron while producing a pair of γ-ray photons that travel away from the annihilation event in opposite directions. If a pair of opposing γ-ray detectors both detect Tray photons within a predetermined period of time, a “coincidence event” is recorded based on the assumption that it was a single annihilation event that occurred along an axis extending between the opposed detectors that produced the detected γ-ray photons.
Conventional PET scanners include arrays of γ-ray detectors that can be provided in a number of configurations including, for example, one or more aligned rings or as one or more pairs of opposed detector plates with lines of response (“LOR”) being defined between opposing pairs of detectors. During operation, the PET scanner collects the radioactivity distribution information within the subject by detecting and accumulating a series of annihilation events originating along each LOR. Regardless of the configuration of the particular PET scanner, the subject will typically be placed at or near the center of the detector array(s) or the near the axis about which detector plates rotate to allow for better sampling rates and provide improved resolution and image quality. Once the position data for a sufficiently large number of annihilation events has been collected, the data may be processed to provide a series of two-dimensional or a three-dimensional image corresponding to the distribution of radiolabeled pharmaceuticals within the subject.
With these conventional designs, the image spatial resolution of a PET system depends on a number of factors including, for example, the intrinsic detector spatial resolution, the acolinearity of the γ-ray photons, and the positron range of the radioisotopes incorporated in the pharmaceutical tracer compounds. Because the acolinearity (or non-colinearity) of the γ-ray photons and the positron range depend on the radioisotope(s) utilized, PET scanner design has tended to focus on improving scanner spatial resolution.
In most conventional PET scanners, a plurality of discrete scintillation crystals coupled to photodetectors have been used to increase the spatial resolution. Improvements in the design of the scintillation crystals and/or photodetectors can improve the spatial resolution, but are limited by an intrinsic spatial resolution that can not be less than one half of the width of the discrete scintillation crystals utilized in the detector. In conventional PET scanners incorporating ring detector geometry, the detector pairs define sampling lines having an average sampling distance that is generally about half of the crystal width. This is in accord with the Nyquist sampling theorem which holds that the smallest object that an imaging system can resolve is twice the size of the sampling distance. In order to achieve image spatial resolution approaching this theoretical limit, various modifications have been incorporated into conventional PET scanners to reduce the sampling distance(s) including, for example, designs in which the detector or the object are moved by a fraction of the detector width and designs that stack discrete crystals in a series of offset layers.
Several additional methods have been developed to improve the resolution of structures within the subject that are below the intrinsic spatial resolution of the detector. One such method includes using a γ-camera coupled to a “pinhole” collimator to produce an effectively enlarged image of the subject and thereby allowing effective resolution of objects smaller than the detector intrinsic spatial resolution. A significant limitation of such designs is the substantial reduction in the number of photons that traverse the “pinhole,” thereby reducing detector efficiency.
Another imaging device is referred to as a “Compton camera” typically includes at least two detectors provided in a cooperative arrangement on a single side of the photon source (subject). The arrangement of the associated detectors provides for the sequential detection of a photon that interacts with at least two of the detectors. The interaction of the photon within the first detector is the result of the Compton effect while the interaction of the photon with the second detector is the result of the photoelectric effect. This sequential detection of a photon by two detectors enables a Compton camera to determine the photon's path without using mechanical collimators such as those conventionally utilized in γ-cameras. In this way Compton cameras tend to exhibit better resolution or discrimination against background than that achieved with conventional γ-cameras, but tend to achieve this at the expense of sensitivity (or fraction of gamma rays detected).
Despite its limitations, PET imaging is an increasingly important medical imaging tool, particularly as the sensitivity and/or resolution are improved and scanners are better able to provide more precise information regarding the nature of the tissues and structures within a scanned subject. Apparatus employed are then used to construct lines of response (LOR) from which images are developed using well known algorithms. In addition to the resolution degrading factors noted above, for example, the range of positrons within the subject and the intrinsic lack of colinearity exhibited by the annihilation γ-rays, PET imaging can compromised by detector spatial and energy resolutions, scattering of γ-rays within the subject before the γ-rays reach the detectors, and scattering in the detectors. In addition, random or accidental coincidences (i.e., not “true” coincidences) will be detected two or more γ-rays from separate annihilation events are detected within an opposed pair of detectors within the predetermined resolving time window. Detection and accumulation of these “false” coincidences limit the statistical precision of image reconstruction.
Events recorded in a typical PET tomograph include 1) the desired true unscattered coincidence events (T), 2) those coincidence events in which one or both of the γ-rays from a single annihilation event are scattered within the subject before entering the detector (TS), and 3) truly random coincidences (R). Adopting this terminology, the T events reflect the desired image information corresponding to the distribution of radioisotopes within the subject while those scattered TS and random R events that are detected as “coincidences” simply contribute “noise” that tends to degrade the resulting image. Although a number of elaborate techniques and algorithms have been developed in an effort to mitigate the effects of the noise on PET images, TS and R noise issues remain and continue to limit the resolution and precision of the resulting PET images.